Method and apparatus for measuring blood volume

ABSTRACT

In one aspect, a conductance catheter is provided for measuring the volume of a fluid. The conductance catheter comprises a series of electrodes and a circuit to compensate for variations in sensitivity of the electrodes in the catheter. In another aspect, a resistivity sensor is provided for determining the resistivity of a fluid. The sensor comprises a series of electrodes spaced such that the total distance between endmost electrodes does not exceed the diameter of the catheter deploying the sensor.

This application claims priority from U.S. Provisional Application No.60/913,734 filed on Apr. 24, 2007, the contents of which areincorporated herein by reference.

TECHNICAL FIELD

The present invention relates generally to data acquisition systems andparticularly to measuring blood volume in a living organism.

DESCRIPTION OF THE PRIOR ART

In the field of cardiac research the standard test for measuring cardiacefficiency is the pressure volume graph. This test correlates LeftVentricle (LV) chamber pressure and volume as the heart contracts andexpands. Pressure and volume values are important for quantifyingefficiency in any pump system, and can be used to calculate volumetricefficiency of such systems. Cardiac efficiency is a useful measurementfor studying heart disease, by quantifying the progress of the diseaseand measuring the effectiveness of the treatment.

Recently, gene altered mice have increased in popularity as a means forstudying heart disease, and for modelling human heart disease.Typically, LV data is measured using a catheter that is inserted intothe LV. The catheter typically has separate instrumentation formeasuring blood pressure and blood volume. There are several drawbacksto using data taken from anaesthetized mice, most significantly the factthat it has been found that cardiovascular data taken from ananaesthetized specimen differs significantly from free-roamingspecimens.

In order to measure cardiovascular data from a free-roaming specimen, aninplanted device is required that can operate while the specimen isactive, and transmit data to the exterior of the specimen forprocessing. This need presents several design problems, notably size andbattery life. Particularly, a reduced size provides a less invasivedevice, and a longer battery life decreases the number of surgicaloperations required to change or recharge a device. The need to reducerepeated trauma due to surgery and the cost of the surgery are drivingreasons for the need to extend battery life in biological implants.These concerns are heightened when extending the application to humanspecimens.

There are numerous devices that have been developed for measuringphysiological pressure in living specimens, e.g., those shown in U.S.Pat. Nos. 4,796,641; 4,846,191; and 6,033,366. These devices include acatheter having a pressure sensor that is inserted into an area in thespecimen having a physiological pressure, such as an artery. The sensorsinclude a pressure transmitting catheter filled with a pressuretransmitting fluid. A pressure transducer communicates with the fluid toprovide an electric pressure signal representing variations inphysiological pressure that can be transmitted to the exterior of thespecimen. These devices are only concerned with measuring pressure, andthe use of a fluid filled catheter can lead to undesirable frequencyresponse characteristics and may exhibit head pressure artefacts.

Other devices, e.g., that shown in U.S. Pat. No. 6,409,674 provide animplantable sensor being anchored to the interior wall of the LV in aliving specimen. The sensor acquires and transmits data from within theheart to an external data receiver. This device is concerned with onlymeasuring a single parameter, and specifically illustrates measuringpressure.

The volume of a liquid in a cylindrical chamber, such as the leftventricle of a heart, can be derived by measuring the conductance of thefluid. The volume (V) can be calculated according to the followingequation:

$V = {\frac{1}{\alpha}\rho \; L^{2}{G.}}$

The variable α represents a non dimensional correction factor attributeto the fact that the electrical field created with catheter basedvolumetry is not typically uniform in its distribution throughout theblood volume. The variable ρ represents the resistivity of the blood inthe LV. It is important for this value to be as accurate as possible andas representative of the actual blood inside the volume being measured.This variable has the potential to change over time or due to researchintervention and thus cannot be considered always consistent. Theresistivity has been shown to vary with temperature, hematocrit andblood velocity. Moreover, it is possible that changes in electrolyteconcentrations also alter resistivity. The variable L represents thedistance between the sensing electrodes, which is fixed by the nature ofthe catheter or other measurement instrument being used. Finally, G isthe actual conductance value that is measured by the electrodes. G mayinclude a correction factor for an overestimation of G caused by theelectric field entering the muscle. Such a correction may then representG as G=G_(block)−G_(correction).

Both α and ρ can be problematic to measure when performing conductancebased cardiac volumetry, since while conductance based cardiac volume isa widely used, its accuracy is limited by the non-uniformity of thetransmitted electric field and the non-fixed value of the bloodresistivity.

Studies have shown that the positioning of a catheter or other measuringdevice in, e.g., a ventricular chamber, is important in accuratelydetermining volume. In particular, this can occur when the catheter isnot centred within the chamber. However, knowing the position of thecatheter within the ventricle can be difficult with existing technologyand relies upon the user's objective skill and experience.

It is therefore an object of the following to obviate or mitigate atleast one of the above-mentioned disadvantages.

SUMMARY OF THE INVENTION

In one aspect, there is provided a sensing tip for measuring the volumeof a fluid comprising one or more electrodes for measuring the volumeand a resistivity sensor disposed on the tip in communication with thefluid to incorporate a current measurement of resistivity duringmeasurement of the volume.

In another aspect, there is provided a system for measuring volume of afluid comprising a sensing tip having a plurality of pairs ofelectrodes, each pair of electrodes being connected to a circuit tocompensate for variations in sensitivity of respective pairs ofelectrodes according to the positioning of the respective pairs alongthe sensing tip.

In yet another aspect, there is provided a method for calibrating asensing tip used for measuring volume of a fluid comprising insertingthe sensing tip into a plurality of cuvettes containing fluids havingdiffering properties, and when in each well, the method comprises:obtaining a plurality of conductance signals using a plurality ofelectrodes on the sensing tip; adjusting the conductance signals tocompensate for variations is sensitivity of the pairs of electrodes dueto the positioning of the electrodes along the sensing tip; obtaining ameasurement of resistivity; and using the conductance signals and themeasurement of resistivity to calibrate the sensing tip.

In yet another aspect, there is provided a method for positioning asensing tip disposed in a ventricle comprising obtaining an excitationwaveform generated by one pair of electrodes disposed on the sensingtip; obtaining a conductance waveform sensed by another pair ofelectrodes disposed on the sensing tip; comparing the waveforms todetermine a phase shift between the waveforms; and adjusting thepositioning until the phase shift is deemed acceptable.

BRIEF DESCRIPTION OF THE DRAWINGS

An embodiment of the invention will now be described by way of exampleonly with reference to the appended drawings wherein:

FIG. 1 pictorially shows a wireless cardiovascular data acquisitionsystem.

FIG. 2 is a schematic representation of the system of FIG. 1.

FIG. 3 is a magnified view of a portion the heart shown in FIG. 1.

FIG. 4 a is a partial plan view of the pressure sensing device of FIG.2.

FIG. 4 b is a sectional view of the sensing device shown in FIG. 4 aalong the line B-B.

FIG. 5 is an electric schematic of the pressure sensing device.

FIG. 6 is a schematic diagram of the transmitter processing module ofFIG. 2.

FIG. 7 is a schematic diagram of the receiver processing module of FIG.2.

FIG. 8 is a timing diagram for the timing controller of FIG. 6.

FIG. 9 is a flow chart showing an acquisition and transmission cycle.

FIG. 10 shows another embodiment of the sensing tip of FIG. 3.

FIG. 11 shows a sensing tip having a resistivity sensor

FIG. 12 shows a circuit for measuring fluid volume.

FIG. 13 shows a calibration circuit.

FIG. 14 shows another calibration circuit.

FIG. 15 illustrates a sectional view of an off-centred catheter in aventricle.

FIG. 16 illustrates a sectional view of a centred catheter in theventricle of FIG. 15.

DETAILED DESCRIPTION OF THE DRAWINGS

A system and method are provided for accurately handling the parametersfor measuring blood volume in real time. In particular, it has beenfound that in order to perform accurate and repeatable conductancemeasurements (parameter G), it is important to eliminate variations inthe measuring electrode sensitivity from the equation. The followingprovides a calibration circuit that takes into account electrodesensitivity at the time of calibration and thus eliminates suchvariability. This also compensates for the non-uniformity of theelectric field generated by a conductance measuring device. In order tocompensate for the electrode sensitivity, an adjustable gain and offsetcan thus be used in each segment of a conductance measurement.

In the following, a resistivity sensor is also provided such that theresistivity (parameter ρ) is measured on a continuous or as-needed basisrather than at time-spaced intervals or relying on a “most recent”measurement. This compensates for variations in resistivity discussedabove and thus provides more accurate volume measurements.

The calibration and resistivity sensor can be incorporated into anydeployable medical device such as an implant or catheter. The followingexample describes an implant transmitter, however, it will beappreciated that the principles for measuring resistivity and forcalibrating volume measurements discussed herein are equally applicableto other devices such as catheters.

It will also be appreciated that the following principles are alsoapplicable to measuring volume in any fluid and should not be consideredlimited only to blood.

Also provided is a system and method for determining the position of acatheter or other measuring device in a ventricle by comparingexcitation and sensed waveforms and correcting the position until thephase angle is minimized.

Referring now to FIG. 1, one embodiment of a wireless cardiovasculardata acquisition system is generally denoted by numeral 10. The system10 operates to measure physical parameters of a heart 12 located withina body 14. The heart 12 and body 14 form part of a living organism, suchas a gene altered mouse or a human. The heart 12 includes a heartchamber, in this example a Left Ventricle (LV) 16 that in partcommunicates with the body 14 via a heart valve 18. A sensing tip 22 issituated in the LV 16 by insertion thereof through the valve 18, and hasa communication path 24 leading to a transmitting device 20 implanted ina portion 15 of the body 14, which in this example is external to theheart 12. In the example shown in FIG. 1, the portion 15 is in proximityof the body's clavicle. It will be appreciated that the transmittingdevice 20 may be situated anywhere as desired, e.g. within the heart 12or heart chamber (i.e. LV 16).

The transmitting device 20 wirelessly transmits data to a receivingdevice 26 that in this example is attached to a belt 27 external to thebody 14. The receiving device 26 may display data on a screen 28 asshown in FIG. 1, and may comprise a keypad 30 for scrolling betweendifferent views. A schematic of the system 10 is shown in FIG. 2.

Referring now to FIG. 2, the path 24 communicates data acquired by thesensing tip 22 to a transmitter processing module 32 in the transmittingdevice 20. The transmitting device 20 is powered by obtaining energyfrom a battery 34, and has a transmitter 36. It will be appreciated thatthe use of a battery 34 is for illustrative purposes only and that anysuitable means for powering the transmitting device 20 may be used suchas power scavenging (converting environmental energy into electricity)or RF power transmission (energy transmitted to the device 20 from anexternal source through a radio frequency signal).

Since the processing module 32 is preferably implanted in the body 14,the signal sent via the transmitter 36 should pass through body tissuebefore reaching the air. The attenuation of an RF signal by differentbody materials is typically highly frequency dependent. Therefore, thetransmitter 36 should be selected so as to minimize the attenuation ofthe signal it transmits. Typically, a lower frequency is preferred totransmit the signals since the lower the frequency, the greater thedepth of penetration. However, the lower the frequency, the higher thewavelength and thus the longer the antenna required at the receivingend. Therefore, the transmitter 36 should be chosen to balance theserequirements depending on the particular application. A suitablefrequency to achieve such a balance is 40 MHz. The power consumed by thetransmitter 36 should also be considered so that it can be faithfullydetected at its receiving end whilst conserving energy.

The transmitting device 20 communicates wirelessly with the receivingdevice 26 through a receiver 40. The device 26 has a receiver processingmodule 38 that is adapted for processing data received from the device20. The device 26 is powered by a battery 42 or suitable AC or DC powersource (not shown). The device 26 has a series of signals (44-50) forproviding electrical representations of measurements acquired using thesensing tip 22, including a pressure signal 44, a volume signal 46, atemperature signal 48, and an electrocardiogram (ECG) signal 50.

In FIG. 2 these signals are shown as being external to the processingmodule 38 and communicably connected to an external computing device 52having an analog-to-digital (A/D) converter 54 connected thereto.However, it will be appreciated that the A/D converter 54 may beincluded in either the processing module 38 or processing module 32, andcomputing device 52 may be replaced by any suitable alternative such asprocessing capabilities provided by the processing module 38. Thecommunicable link between the receiving device 26 and the computingdevice 52 and/or A/D converter 54 may be any hardwired or wirelesscommunication channel, e.g., using Bluetooth technology.

The computing device 52, external or internal to the receiving device26, may be any device that is capable of acquiring data andcommunicating with the processing module 38. In the example shown inFIG. 2, the device 52 is a standard personal computer (PC) having amonitor, central processing unit (CPU), keyboard, and mouse.

The sensing tip 22 is shown in greater detail in FIG. 3. The sensing tip22 has a rounded end 70 to facilitate the deployment thereof through thevalve 18. In this example, a proximal electrode 62 and a distalelectrode 60 each following the circumference of the sensing tip 22flank a pair of inner electrodes 64, 66, a pressure sensing device 68,and a temperature sensing device 69. The electrodes 60, 62, 64 and 66are used to measure the volume of blood in the LV 16 and are hereincollectively referred to as the volume sensing device denoted by numeral67. The proximal electrode 62 transmits a signal, and the distalelectrode receives same to create an electric field in the LV 16. Theinner electrodes 64, 66 sense this electric field to perform aconductance measurement indicative of the volume in the LV 16. The innerelectrodes 64, 66 can be modeled conceptually as measurement probes oneither side of a “resistor”, wherein the “resistor” represents theresistivity of the blood in the LV 16, the inner electrodes 64, arearranged to measure the potential across the “resistor”. The volumemeasurement and/or volume signal may also be referred to as aconductance measurement and/or conductance signal respectively, and itwill be appreciated that this terminology may herein be consideredinterchangeable.

The pressure sensing device 68 is used to sense the pressure of theblood in the LV The temperature sensing device 69 is used to sense thetemperature of the body 14, since it is substantially uniformthroughout. The temperature sensing device 69 is preferably comprised ofa thermistor or equivalent component. The volume sensing device 67,pressure sensing device 68, and temperature sensing device 69communicate data to the transmitting device 20 through the path 24, thusthe path 24 typically carries a number of wires, enabling data to betransmitted from the sensing tip 22 to the device 20. The length of thepath 24 is dependent upon the location of the device 20 relative to theheart 12.

Although the temperature sensing device 69 is shown in FIG. 3 as part ofthe sensing tip 22, it will be appreciated that the device 69 may besituated anywhere in the body 14 enabling the internal temperature ofthe body 14 to be measured, and this may be inside or outside of theheart 12.

An embodiment of the sensing tip 22 is shown in FIGS. 4 a and 4 b. Itwill be appreciated that the relative dimensions of the sensing tip 22have been exaggerated for illustrative purposes only. The pressuresensing device 68 may be any device capable of sensing a pressure. Inthis example, the pressure sensing device comprises a piezoresistivedeflection sensor, specifically a cantilevered sensor beam 80 having abase portion 82 that is attached to the housing of the sensing tip 22. Abase window 85 in the sensing tip 22 enables the base of the beam 80 toexperience external pressure, and a tip window 86 enables the tip of thebeam 80 to experience external pressure. A layer of sealant 88 inhibitsthe beam 80 from direct contact with its surrounding environment.However, the layer 88 permits external pressure to effect flexure of thebeam 80 due to variations in the pressure of the surrounding blood. Itcan be seen in FIG. 4 b that electrical wires run from the sensingdevices 67, 68 and 69 to the path 24.

An implementation of the beam 80 is shown schematically in FIG. 5, beinga strain gauge sensor, on which two resistors R_(x1) and R_(x2) aremounted. When the beam bends as a result of a pressure experiencedthereby, the resistances of these resistors change in oppositedirections. That is, the resistance of one of the resistors increaseswhile that of the other one decreases. As a result, the accompanyingelectronic circuits may be designed in a fully differential architecturewhich provides a higher signal to noise ratio (SNR) compared to a singleended architecture.

The following lists suitable specifications for the pressure sensingdevice 68, but shall in no way be considered limited thereto: nominalresistance of each resistor R_(x1), R_(x2) being 10,000 Ohms; gaugefactor of 70-80; total resistor manufacturing tolerance of +/−10-15%;maximum resistance value mismatch between the resistors of 2.4%;temperature coefficient of resistance of +5%/100° F.; and a breakdownvoltage of 20V.

These exemplary specifications illustrate that typically there may benon-idealities for the sensing device 68 that would preferably beaddressed when designing the circuitry therefor. For instance, due toprocess variations, the resistances of R_(x1) and R_(x2) are in alllikelihood not going to be equal. This may generate some offset at theoutput. Moreover, since the resistance of the resistors R_(x1) andR_(x2) is a temperature dependent parameter, the temperature coefficientof resistance (TCR) may cause an offset due to mismatch. Hence, even ifthe offset is cancelled at one temperature it may not be zero at anothertemperature. Finally, the temperature coefficient of the gauge factor(TCGF) makes the gain of the sensing device 68, temperature dependent.

The above parameters are typically sources for measurement inaccuracies.As a result, the output of the sensing device 68 may have some offseterror and be dependent on temperature. In order to compensate for theabove parameters, typically a signal conditioning scheme is utilized. Inthe example shown in FIG. 5, a Wheatstone bridge configuration is usedto measure the resistance variations with two current sources I₁ and I₂.

As indicated above, R_(x1) and R_(x2) change in opposite direction as afunction of strain or equivalently blood pressure in the heart as:R_(x1)=R₀₁(1+GF.x) and R_(x2)=R₀₂(1+GF.x) where R₀₁ and R₀₂ are thesensor resistances at zero strain, GF is the gauge factor of the sensingdevice 68, and x is the strain. The two current sources I₁ and I₂complete the bridge, and are preferably integrated into the processingmodule 32 as shown in FIG. 5. In order to cancel out the resistormismatch, TCR, and TCGF, the following equations should be valid:R₀₁I₀₂−R₀₁I₀₁=0; and TCI=(TCR+TCGF); where TCI represents thetemperature coefficient of the current sources, R₀₁ and R₀₂ representthe resistor values at the reference temperature, and I₀₁ and I₀₂represent the current of the two current sources at the referencetemperature. The technology used to implement the processing module 32should be capable of implementing a current source with any specifictemperature coefficient, and the current sources should preferably bedesigned to have the lowest possible supply voltage sensitivity.

A block diagram of the transmitter processing module 32 is shown in FIG.6. The module 32 comprises a sensing block 90 and a transmitting block92 controlled by a timing controller 94. The battery 34 which isconnected to the module 32 may be controlled by a switch The battery 34is preferably a miniature battery of a suitable size and having abattery life that is as long as possible. A suitable battery has a lifeof 180 mAh, weight of 2.3 g, 1.5 Vdc, and a volume of 0.57 cc. Theswitch 96 may be, e.g., magnetic or radio controlled, i.e. any suitabledevice capable of controlling the main power to the module 32 from thebattery 34. Between the timing controller 94 and the switch 96 is avoltage regulator that provides a regulated voltage to the timingcontroller 94 for controlling the blocks 90 and 92. With the abovebattery specifications, a suitable regulated voltage is a 1V output.

The sensing block 90 includes a current source block 100 for thepressure sensing device 68 (described above with current sources I₁ andI₂) to compensate for sensor non-idealities, and are the basis oftemperature compensation for the pressure sensing device 68. The block90 also includes a conductance current source 102 for generating theelectric field using the electrodes 60 and 62; and a thermistor currentsupply 104 for the temperature sensing device 69, that preferablycomprises a high resistance thermistor for minimal current drain. Theoutputs from these current sources (100-104) are sent to the sensing tip22 over the path 24.

The measurements acquired by the sensing devices 67, 68 and 69 are sentback to the sensing block 90 over the path 24. The temperature signal isfed through an amplifier 106 and sampled and held for transmission by asample and hold component 112. Similarly, the pressure signal is fed toan amplifier 110 and sample and hold component 116; and the volumesignal is fed to an amplifier 108 and sample and hold component 114. Theamplifiers 106, 108 and 110 are preferably used to encourage thefidelity of the signals. The sample and hold components 112, 114 and 116hold the signal samples while the timing controller 94 switches powerfrom the sensing block 90 to the transmission block 92.

The transmission block 92 has a multiplexer 118 and a voltage controlledoscillator (VCO) 120. The multiplexer 118 will read the samples from theblocks 112-116 and arrange the signals for transmission by the VCO 120.For example, the multiplexer 118 may arrange the signals in sequentialorder for transmission. The VCO 120 is connected to an antenna 121 andtogether make up the transmitter 36 shown in FIG. 2. A suitable VCO 120is a Colpitts type that consumes an average current of 32 μA. Theantenna 121 is preferably connected in parallel with the frequencydetermining inductor of the VCO 120, and preferably serves as an FMtransmitter with a 42 MHz transmission frequency.

A block diagram of the receiver processing module 38 is shown in FIG. 7.The module 38 comprises a demultiplexer 122 connected to the receiver 40of the receiving device 26. The demultiplexer 122 separates the signalsthat have been transmitted by the transmitter 36 and received by thereceiver 40. If the signals are transmitted as analog signals, thedemultiplexer 122 separates the received signal into individual analogsignals, and in this example would provide three individual signals, atemperature signal 124, a pressure signal 126, and a volume signal 128.The temperature signal 124 may be immediately available as output 48,and the pressure signal 126 may be immediately available as output 44for further processing and/or transmission to the computing device 52.It will be appreciated that the module 38 may also comprise a furtherinternal component for processing and analysing the signals 124, 126 and128, e.g., for display purposes. Moreover, the module 38 may comprise analarm or other device to notify a wearer of the receiving device 26 ofabnormal heart conditions. The display 28 may also be used with suchadditional processing to output heart parameters or a computed indexthat represents heart health.

The volume signal 128 may be sent through a buffer 129 and be availableas output 46. The volume signal 128 may also be captured at block 130for further processing to extract the ECG signal. This preliminarysignal 130 is preferably converted using an analog-to-digital converter(A/D) 132, which enables signal manipulation while preserving theintegrity of the original signal. It will be appreciated that the A/D132 would not be needed if the signals received have already beenconverted to digital signals. The A/D 132 has two identical outputs, oneof which is input to a digital signal processor (DSP) 134. The DSP 134is used to clean the ECG signal from the volume signal, and allows forcomplex signal processing. The extraction of the ECG signal is describedin greater detail later.

The signal emerging from the DSP 134 is inverted by an inverter 136. Theinverter 136 may also be part of the DSP 134. The other output from theA/D 132 is buffered by the buffer 138 and the inverted signal and thebuffered signal are summed at 140 to produce the ECG signal 142 that mayalso be available as output 45. The buffer 138 is used to maintain thesynchronicity of the raw volume signal and the digitally manipulatedversion (i.e. by the DSP 134). The delay imposed by the DSP 134 wouldotherwise affect the results of the sum 140. The summer 140 adds the twovolume signals, and since one has been inverted, the conductance part ofthe volume signal will be eliminated and the remaining signal willrepresent the ECG signal 142.

The sensing block 90 and the transmitting block 92 are selectivelypowered using the timing controller 94 in order to conserve power. Atiming diagram is shown in FIG. 8 illustrating the operation of thetiming controller 94. The period T represents an entire monitoring cyclefor the system 10 including measurement and transmission. Specifically,T₁ represents the period in which the sensing block 90 is powered inorder to obtain the necessary measurements and sample and hold thesignals; and T₂ represents the period in which the transmitting block 92is powered in order to execute transmission of data from thetransmitting device 20 to the receiving device 26.

For example, a 2 kHz sampling rate provides a period T of 500 μs tosample and transmit data. If the acquisition period T₂ is 20 μs, andtransmission period T₃ is 50 μs, there exists 430 μs during each cycle,in which either the block 90 or the block 92 is waiting. The timingcontroller 94 uses this timing scheme to selectively turn off either theblock 90 or block 92 that is not being used to conserver power, whichprovides an increase in battery life.

Another benefit arises from using such an energy saving timing scheme,namely the reduction of noise. Specifically, since the block 90 ispowered whilst the block 92 is not, the transmitter 36 will not beaffected by the noise generated by the signal conditioning, and,conversely, the sensing circuitry (block 90) will not be subject tonoise from the transmitter 36. A 10 μs period, represented by T₃, isleft between the end of one period and the beginning of the next, whichenables any circuitry that needs stabilizing to do so.

Therefore, since the transmitting block 92 typically cannot transmitdata that has not yet been collected, it would be wasting power whilethe sensing block 90 is performs its function. If the transmitting block92 is turned off when it is not needed, power is not consumed, and thusconserved. Similarly, the sensing block 90 typically is not adding anydata while the transmitter 36 is sending the previous sample, and thusdoes not need to consume power during that time.

FIG. 9 shows a flow chart illustrating an example of the steps taken bythe system 10 during one complete cycle T, and the subsequent processingby the receiving device 26. The sensing block 90 is powered whichenables the current sources to power the measurement devices 67, 68 and69 and obtain the measurements. These measurements are then amplifiedand undergo a sample and hold. The sensing block 90 is then powered“off” and the transmitting block 92 is powered “on”, wherein the timelag between theses steps is represented by T₃ as explained above. Oncethe block 92 has power, the multiplexer 118 is then able to obtain thesignals stored in the sample and hold components 112-116, and combinethese signals for transmission. In this example the multiplexer 118preferably operates by arranging the signals in a particular sequentialorder that would be known to the demultiplexer 122 in order to enablethe demultiplexer 122 to separate the signals at the receiving end.

The multiplexer 18 passes this “combined” signal to the VCO 120 thatuses the antenna 121 to transmit the “combined” signal to receivingdevice 26. At this point, a complete measurement cycle has beenexecuted, and the signal that has been transmitted continues to thereceiving device 26 for further processing and/or output. Thetransmitting device 20 may then repeat this cycle as required ordesired.

The receiving device 26 receives the “combined” signal from the receiver40. The signal is passed to the demultiplexer 122 where it is separatedinto its components. The temperature and pressure signals 124 and 126respectively, may be available as outputs or for further processing bythe module 38. The volume signal 128 may be buffered and output at 46,and may also be obtained for extracting the ECG signal 142 and providingoutput 45. The extraction of the ECG signal 142 from the raw volumesignal 128 is described in greater detail below, while referring to thefunctional blocks shown in FIG. 7 that relate thereto.

As indicated above, the conductance or volume signal 128 acquired usingthe volume sensing device 67 is used to extract the ECG signal 142.

The conductance signal acquired using the volume electrodes 67 consistsof the conductance value of the blood in the LV 16, any noise generatedby the system or in the environment, and the ECG signal 142 that ispicked up as a component of environmental noise. As described above, inthis example, the raw signals are collected and transmitted, e.g., as acombined analog waveform, without performing any signal conditioning, tothe receiving device When the combined signal is received by thereceiving device 26, the individual pressure, volume and temperaturesignals (124, 126 and 128) are separated, and a process begins toseparate the various components of the volume signal 128 (i.e. at 130).

The conductance signal 128 is the result of an electrical fieldgenerated, by means of the electrodes 60, 62, from the apex of the heartto the carotid artery. Due to myocardial contact of the conductancerings, the resulting conductance signal will also carry the ECG signal.It is generally common practice to use signal conditioning and filteringto eliminate the environmental and ECG noise components to extract theconductance signal 128. In this embodiment, signal conditioning is usedto not only remove the ECG component of noise to extract the conductancesignal, but also to separately condition the ECG signal 142 to removethe conductance portion of the signal. The result is that an ECG signal142 can be collected without introducing any additional instrumentationinto the LV 16. Therefore, the sensing tip 22 can be used to provide amore thorough cardiac assessment, using a single device.

Once the signal is obtained at 130, an A/D converter 132 in theprocessing module 38 converts the raw signal to a digital signal andpasses the signal to each of an ECG digital signal processor (DSP) and abuffer 138. Once the respective signals are processed, they are summedand a final ECG signal 142 is produced.

In another embodiment, the volume sensing device 67 comprises aplurality of inner electrode rings, for example four as shown in FIG.10. Since the optimal conductance measurement is performed bytransmitting along the entire length of the LV 16, and differentorganisms have different sized hearts 12, it may be desirable toincorporate multiple sets of inner electrode ring pairs. In FIG. 10, theLV 16 shown in FIG. 3 is provided, as well as an LV 1016 from a smallerorganism shown in dashed lines. The pair 164, 166 is similar to the pair64, described above, however, the sensing tip 22 now includes the pairs168, 170; 172, 174; and 176, 178 arranged progressively closer togetherand situated between the outer electrode pair 60, 62.

In such an embodiment, it may be possible to selectively operate any ofthe electrode rings as a transmitting ring, but typically the electrode60 would remain as the receiving electrode. In the example shown in FIG.10, the electrode 170 would be selected as the optimal transmittingelectrode for the LV 1016 and then the inner sensing electrode pairwould comprise the electrodes 164 and 174. Therefore, numerousconfigurations of receiving, and sensing electrodes can be selectivelychosen in order to obtain an optimal conductance signal, depending onthe size of the LV (e.g. 16 or 1016).

Therefore, the system 10 enables the monitoring of a heart in a livingorganism by measuring both pressure and volume in a chamber of theheart, preferably the LV 16. The pressure and volume measurements areacquired using a single sensing tip 22 and are communicated to atransmitting device 20 to be wirelessly transmitted to a receivingdevice 26, wherein they are used to monitor the heart. The system 10 mayalso incorporate a temperature measurement that can be transmitted withthe volume and pressure measurement to provide further data formonitoring. The system 10 may also extract an ECG signal from the volumemeasurement. This allows the monitoring of up to four signals that canbe used to determine the health of a heart.

In addition to a compact design, the system 10 may also incorporate anenergy saving timing scheme that reduces the power required peracquisition cycle and thus increases the operational lifetime of thetransmitting device 20.

As noted above a calibration scheme and resistivity meter can also beincorporated into a system such as system 10. In yet another embodiment,shown in FIG. 11, catheter 22 includes a resistivity sensor (generallynumeral 200). In one implementation, resistivity sensor 200 a comprisesa series electrodes 202 arranged substantially parallel to the axis ofthe catheter 22. In another implementation, resistivity meter 200 bcomprises a set of spaced apart rings 204, e.g. four. The spacing issuch that D is less than the diameter of the rings 204. The resultingfield is thus small enough that it does not experience any significanteffect from changing volume. In general, the resistivity sensor 200 canbe configured using any four electrodes spaced such that the distanceapart of the further electrodes does not exceed the diameter of thecatheter.

The resistivity sensor 200 operates on a similar principle to that of aconductance catheter. Four electrodes are deployed in series, and aconstant current flows through the fluid. The current travels betweenthe two outermost electrodes 204. The two inner electrodes are used tosense the voltage created by the resistance of the blood. By measuringvoltage and by applying a known current, the resistance can be obtainedaccording to the relationship V=IR.

Using the configuration shown in FIG. 11, a continuous or as-neededresistivity measurement can be made, which in turn enables more accurateblood volume measurements.

As noted above, the resistivity sensor 200 can be used to measure thevolume of any fluid, e.g. by considering the equation for a volume offluid in a cylinder, namely V=ρL²G. This equation is a simplifiedversion of the typical equation used to measure LV volume, namely whereα=1 and disregarding any correction factor for G. It has been found thatif ρ changes in a cylindrical volume, the volume reading changes by adirectly proportional amount, even though the volume in the cylinder isactually the same. By using the resistivity sensor 200, the value for ρcan be changed at any time and to any desired degree, while not changingthe value for the volume of the fluid. This enables the system to becalibrated using a known volume, independent of the value for ρ.

It can therefore be seen that by incorporating a resistivity sensor 200,an accurate and current resistivity value can be obtained in real timerather than relying on the accuracy of the latest measurement. Also,where resistivity measurements are obtained by drawing blood,significant discomfort can be avoided. Although two different variationsof sensor 200 are shown, it will be appreciated that typically only onevariation would be used.

A schematic diagram of a volume measurement circuit 206 is shown in FIG.12. The circuit 206 provides a comprehensive system for calibratingconductance signals for estimating blood volume, e.g. in an LV. Theelectric field distribution from a dipole catheter is a series ofconcentric rings emanating from the distal and proximal rings, asdiscussed above. Due to this configuration, the voltage sensed bydifferent ring segments 208 varies depending on where they are locatedin the field. The effect is similar to a line of listeners in front of aspeaker: those is the front will hear a louder signal than those in therear.

In the circuit 206 shown in FIG. 12, each segment 208 of the catheter 22is assigned an individual gain and offset circuit 210. This enables acustom gain to be applied to each pair of rings. The effect of adjustingthe gain is beneficial for both signals that may be too “weak” as wellas those that could be too “strong”. In either case, interpretation canbe affected and thus compensating for such effects improves the qualityof the measurement. As can be seen in FIG. 12, the segments are summedat stage 212 and a display 214, e.g. a PC monitor is used to view thevolume measurement.

In order to set the individual gains, a calibration fluid having a knownconductance value can be used, which corresponds to the fluid beingmeasured, (blood or otherwise). This allows the signals to benormalized. In general, the catheter 22 is placed in a well containing afluid of a known conductance value. The readings can then be displayedon an electronic display. The circuit then adjusts the gains for theindividual segments such that the signals are all reading the samevoltage output for a given solution. In addition to adjusting voltageamplitude, the circuit 206 can also adjust for baseline voltage suchthat all the signals have the same span and output for a given segmentalvolume. The calibration fluid can also be used to calibrate the system10 for linear output of volume. By using a series of graduate volumes,the output of the catheter 22 can be recorded as it moves from onevolume to the next.

The calibration system can also use a series of wells into which isplaced fluids of differing but know conductivities. This enablesverification of the linearity of the system 10 as it measures the valueof conductance G for each solution. The voltage outputs can be displayedin a plot to verify linearity and hence accuracy of the system.

FIGS. 13 and 14 illustrate two arrangements for calibrating the catheter22 to account for the actual and varying sensitivity of the electrodes(60-66).

In the arrangement shown in FIG. 13, two wells of known volume in avolume cuvette 220 are used to determine the volume calibration slope.In the example shown in FIG. 13, a multi-segment control box 222connected to the catheter 22 provides a segmental output that isproportional to the volume of liquid defined by the diameter of thecylinder in the cuvette 220, and the spacing of each of the segments208. As noted above, the ρ sensor 200 incorporated into the catheter 22can be used to compensate for any chance in ρ that may occur. Bydetermining the value of ρ and incorporating such a value into thecalibration measurement, it is possible to change the ρ value of thefluid being measured, without changing the output voltage. The volumemeasured is then independent of ρ. For this arrangement, the actualvalue of conductivity is not required. It can be seen that in thisarrangement, a correction factor can be maintained if it should change.

The control box 222 also includes the gain and offset circuitry 210,which is adjusted such that all the segments 208 deliver the samevoltage when in a given cuvette well. The gain/offset circuit 210 can beapplied in the same way regardless of the method of calibration. Forexample, using 20 and 50 ml volume wells, when placed in the 20 mlvolume well, the following readings may be observed, shown in Table 1below:

TABLE 1 Example calibration readings V @ V @ Gain/Offset Gain/OffsetSegment 20 ml 50 ml correction @ 20 ml correction @ 50 ml 1 2 6 −2 0 2 03 −2 0 3 5 8 −2 0 4 3 5 −2 0 5 −1  −2  −2 0

One concern is that with a +/−10 volt A/D 224, the range could beexceeded if the properties of the fluid change, e.g. when a saline bolusin injected to correct for parallel conductance caused by myocardialcontribution. The system described herein enables the user to dial inthe span and gain desired by the user and being suitable to a wide rangeof subjects. The ρ sensor 200 in this arrangement would not need to becalibrated as it would not see a volume change going from one cuvettewell to another. The sensitivity of the ρ sensor 200 can be correctedhowever if it is not reading the correct values for the fluid. This canbe done since in the arrangement shown, the ρ values are known for eachfluid. Being able to account for varying electrode sensitivities can bebeneficial for the accuracy of the measurements.

The output from the control box 222 is fed into an A/D converter 224 andthe volume is calibrated according to the linear relationship y=mx+b.Two of the cuvette wells, with known volumes, can thus be used todetermine the value for m. In this arrangement, the value of ρ is beingcompensated for in the control box 222. It is also possible to outputthe value of ρ to the PC 214 and use it as a scaling factor for thevolume calibration.

In the arrangement shown in FIG. 14, an equal volume cuvette 230 can beused, which has a plurality of wells with equal volumes. To calibrate, afluid having a known and precise value of conductance can be insertedinto each well. In this arrangement, a control box 232 provides asegmental output that is proportional to the conductance of the liquiddefined by the diameter of the cylinder (fluid well) and the spacing ofeach segment 208. Both the conductance and resistivity voltagesgenerated by the calibration fluids are sent from the control box to thePC 214, with the segmental output again being fed through A/D converter224. The PC 214 determines the slope for both the conductance G and theresistivity ρ and the values for each segment are summed to provideG_(total).

The values of ρ and G are applied to the volume formula

$V = {\frac{1}{\alpha}\rho \; L^{2}{G.}}$

Since both ρ and G have been obtained in a controlled and accuratemethod and can be monitored on an ongoing basis, the total volumeaccuracy can be maintained.

It will be appreciated that the calibration of the circuit 206 can bedone prior to use of the catheter 22, but may also be configured to beperformed periodically while the catheter 22 is deployed to enable realtime calibration. In this way, a known voltage can be used to representa conductance level.

It has also been found that the position of a catheter or othermeasuring device can be determined by comparing the excitation andmeasured conductance waveforms using a phase angle detector. Bycomparing such waveforms, it is possible to observe a phase shiftbetween the two waveforms, which is caused by the capacitive nature ofthe myocardium. This is caused by the catheter being off-centred in theventricle. An AC waveform moving through a capacitor will incur a phasedelay. Since blood is a purely resistive material, it does not addappreciably to the observed phase shift. As will be discussed below, ithas been recognized that by superimposing the sinusoidal electrodestimulation waveform with the sinusoidal sensed voltage waveform, thecatheter electrode position within the ventricle can be determined.

The above can be incorporated into a real-time feedback scheme thatallows the use of such a measurement of phase-angle to adjust thecatheter electrode position so that they are in the optimum positionwithin the ventricle. Optimum positioning is obtained when the phaseangle is at a minimum, indicating that the electric field is in a centerposition that minimizes field incursion into the myocardium.

Turning now to FIGS. 15 and 16, a sectional view of a ventricle 16 isshown. In FIG. 15, the sensing tip 22 of a catheter is offset oroff-centred in the ventricle 16 and thus it can be seen that theexcitation signal 202 is affected by the capacitive effect of themyocardium The sensed voltage 304 is thus offset from the excitationsignal 302 thus creating a phase angle 306 when the two waveforms areobserved. It will be appreciated that the waveforms 302, 304 can beviewed using any monitoring equipment, e.g. in real-time measurements,during calibration stages etc. FIG. 16 shows that by centering thesensing tip 22 in the ventricle 16, the excitation signal 302 travelsubstantially through the blood rather than the capacitive myocardium300 and thus the phase angle 306 is minimized. Based on thisobservation, the waveforms 302, 304 can be superimposed in a display asshown in FIGS. 15 and 16 (e.g. using display 214 or terminal 52) toprovide feedback to a user so that the positioning of the sensing tip 22can be adjusted until the phase angle 306 is minimized thus indicatingwhen the sensing tip 22 is substantially centered. In anotherembodiment, the zero-crossings of the waveforms can be compared toprovide a numerical offset factor that can be observed as the sensingtip 22 is adjusted until this offset value is minimized. By providingsuch feedback, further accuracy of the volume measurement can beattained.

Although the invention has been described with reference to certainspecific embodiments, various modifications thereof will be apparent tothose skilled in the art.

1. A sensing tip for measuring the volume of a fluid comprising one ormore electrodes for measuring said volume and a resistivity sensordisposed on said tip in communication with said fluid to incorporate acurrent measurement of resistivity during measurement of said volume. 2.The sensing tip according to claim 1 wherein said resistivity sensorcomprises a series of resistivity electrodes spaced from each other suchthat a total distance between endmost ones of said resistivityelectrodes does not exceed the diameter of said sensing tip.
 3. Thesensing tip according to claim 2 wherein said resistivity electrodes arediametrically spaced.
 4. The sensing tip according to claim 2 whereinsaid resistivity electrodes are axially spaced.
 5. The sensing tipaccording to claim 1 configured for continuously obtaining saidmeasurement of resistivity.
 6. A system for measuring volume of a fluidcomprising a sensing tip having a plurality of pairs of electrodes, eachpair of electrodes being connected to a circuit to compensate forvariations in sensitivity of respective pairs of electrodes according tothe positioning of said respective pairs along said sensing tip.
 7. Thesystem according to claim 6 wherein said circuit comprises individualgain and offset circuits to adjust the strength of a respective signal.8. The system according to claim 7 comprising an output for displayingsaid respective signal to facilitate adjustment of said gain and offsetcircuits.
 9. The system according to claim 6 comprising a control boxfor calibrating said sensing tip, said control box including a gain andoffset circuit for each pair of electrodes to adjust the strength of arespective signal; and a monitor for displaying said respective signalto facilitate adjustment of said gain and offset circuits.
 10. Thesystem according to claim 9 wherein said control box obtains aresistivity measurement from said sensing tip and incorporates a valueof resistivity into a calibration measurement.
 11. The system accordingto claim 9 comprising a plurality of cuvettes for containing fluids ofdiffering properties.
 12. A method for calibrating a sensing tip usedfor measuring volume of a fluid comprising inserting said sensing tipinto a plurality of cuvettes containing fluids having differingproperties, and when in each well, said method comprises: obtaining aplurality of conductance signals using a plurality of electrodes on saidsensing tip; adjusting said conductance signals to compensate forvariations is sensitivity of said pairs of electrodes due to thepositioning of said electrodes along said sensing tip; obtaining ameasurement of resistivity; and using said conductance signals and saidmeasurement of resistivity to calibrate said sensing tip.
 13. The methodaccording to claim 12 wherein said differing properties comprisesdiffering volumes and said measurement of resistivity is obtained usinga resistivity measurement on said sensing tip.
 14. The method accordingto claim 12 wherein volume measurements obtained by said sensing tip arecalibrated according to a linear relationship.
 15. The method accordingto claim 12 wherein said differing properties comprises differing knownvalues of conductance and said method comprises providing a segmentaloutput that is proportional to the conductance of the fluid in therespective cuvette.
 16. A method for positioning a sensing tip disposedin a ventricle comprising obtaining an excitation waveform generated byone pair of electrodes disposed on said sensing tip; obtaining aconductance waveform sensed by another pair of electrodes disposed onsaid sensing tip; comparing said waveforms to determine a phase shiftbetween said waveforms; and adjusting said positioning until said phaseshift is deemed acceptable.
 17. The method according to claim 16 whereinsaid comparing comprises superimposing said waveforms to provide avisual indication of said phase shift.
 18. The method according to claim16 wherein said phase shift is determined by comparing zero-crossings ofsaid waveforms.
 19. The method according to claim 16 performed in realtime while said sensing tip is measuring volume of fluid in saidventricle.
 20. The method according to claim 16 performed during acalibration operation.